1. Field of the Invention
The present invention relates to a programmable auditory prosthesis, such as a hearing aid or cochlear implant. In particular, the invention is an auditory prosthesis that adjusts its sound processing characteristics in a particular acoustic environment in a manner that is similar or identical to that previously determined by the user of the prosthesis as optimal for that environment.
2. Related Art
Hearing loss, which may be due to many different causes, is generally of two types, conductive and sensorineural. Of these types, conductive hearing loss occurs where the normal mechanical pathways for sound to reach the hair cells in the cochlea are impeded, for example, by damage to the ossicles. Conductive hearing loss may often be helped by use of conventional hearing aid systems, which comprise a microphone, an amplifier and a receiver (miniature speaker) for amplifying detected sounds so that acoustic information does reach the cochlea and the hair cells. Since the elevation of the minimum detectable sound pressure level may vary with the frequency of an acoustic test stimulus, the amplifier may be preceded by, or comprise of, a bank of filters to enable different frequency components of the signal to be amplified by different amounts.
Sensorineural hearing loss occurs where the hair cells in the cochlea and the attached auditory nerve fibres are damaged or destroyed. Sensorineural hearing loss results in an increase in the minimum detectable sound pressure level, which often varies with the frequency of the test stimulus. However, in contrast to conductive hearing loss, the sound pressure level that is uncomfortably loud at a given test frequency is often approximately the same as for people with normal hearing. The result is a reduction in the dynamic range of sound pressure levels that are audible yet not uncomfortably loud with the impaired ear, and this dynamic range of the impaired ear may vary considerably with the frequency of the acoustic test stimulus. For this reason, sensorineural hearing loss is often treated with hearing aid systems that employ non-linear amplification to compress the dynamic range of common sounds towards the dynamic range of the impaired ear. Such systems may use a filter bank that is followed by a bank of compressive amplifiers, so that the dynamic range of the signal is reduced by an amount that is considered appropriate for the dynamic range of the impaired ear in each band.
In many people who are profoundly deaf, the reason for deafness is absence of, or destruction of, the hair cells in the cochlea which transduce acoustic signals into nerve impulses. These people are thus unable to derive suitable benefit from hearing aid systems, no matter how much the acoustic stimulus is amplified, because there is damage to or absence of the mechanism for nerve impulses to be generated from sound in the normal manner. It is for this purpose that cochlear implant systems have been developed. Such systems bypass the hair cells in the cochlea and directly deliver electrical stimulation to the auditory nerve fibres, thereby allowing the brain to perceive a hearing sensation resembling the natural hearing sensation normally delivered to the auditory nerve. U.S. Pat. No. 4,532,930, the contents of which are incorporated herein by reference, provides a description of one type of traditional cochlear implant system.
Cochlear implant systems have typically consisted of two essential components, an external component commonly referred to as a processor unit and an internal implanted component commonly referred to as a stimulator/receiver unit. Traditionally, both of these components have cooperated together to provide the sound sensation to a user.
The external component has traditionally consisted of a microphone for detecting sounds, such as speech and environmental sounds, a speech processor that converts the detected sounds, particularly speech, into a coded signal, a power source such as a battery, and an external transmitter coil.
The coded signal output by the speech processor is transmitted transcutaneously to the implanted stimulator/receiver unit situated within a recess of the temporal bone of the user. This transcutaneous transmission occurs via the external transmitter coil which is positioned to communicate with an implanted receiver coil provided with the stimulator/receiver unit. This communication serves two essential purposes, firstly to transcutaneously transmit the coded sound signal and secondly to provide power to the implanted stimulator/receiver unit. Conventionally, this link has been in the form of an RF link, but other such links have been proposed and implemented with varying degrees of success.
The implanted stimulator/receiver unit traditionally includes a receiver coil that receives the coded signal and power from the external processor component, and a stimulator that processes the coded signal and outputs a stimulation signal to an intracochlea electrode assembly which applies the electrical stimulation directly to the auditory nerve producing a hearing sensation corresponding to the original detected sound.
Different users of auditory prostheses require differing outputs from their prosthesis to suit their individual requirements. This is the case even when individual users may clinically be regarded as having identical hearing loss profiles, are utilising identical prostheses, and when exposed to essentially identical acoustic environments. Because of this, sound processing schemes for hearing aids and cochlear implants typically contain a number of parameters for which the values can be adjusted to suit the requirements of individual users. Examples of such parameters include the sensitivity to incoming sounds and the variation of the frequency response. Typically, the parameter values are selected either by the prosthesis user in everyday situations (eg. the sensitivity (volume), frequency response), or by the clinician at the time the prosthesis is fitted (eg. the baseline frequency response and the rate at which the frequency response and sensitivity vary as the input level varies).
In more recent times, there has been a trend to provide auditory prostheses with an increasing number of adjustable parameters that can or must be adjusted to optimise performance. This increase has, however, highlighted the problem that there does not always exist a reliable prescriptive method for selecting the optimum values for the individual user, particularly as some optimum values may vary among individuals who have hearing loss profiles that may clinically be regarded as identical. It is accordingly often necessary for the clinician to make adjustments to the prosthesis based on the user's reported experiences away from the clinic, and the need to return to the clinic for these adjustments can be time consuming and inefficient.
One example of a hearing aid that can receive the impressions of a hearing aid user and take these into consideration during operation is described in U.S. Pat. No. 5,604,812. This document describes a hearing aid that has a memory that can store so-called “unsharp inputs”, or impressions of the hearing aid wearer about prevailing ambient conditions and/or the volume of the output of the hearing aid, prescribable algorithms, hearing loss data and hearing aid characteristic data. A fuzzy logic module uses this data and control signals from an input signal analysis unit to calculate the output setting parameters for the hearing aid. The behaviour rules of the fuzzy logic module and/or the prescribable algorithms are calculated from the data stored in the aid's memory by a neural network, with the neural network being typically implemented on a personal computer that is connected to the aid. A problem with such an aid is the complexity of the processing scheme and an undesirably large drain on the aid's power source, typically an on-board battery. Further to this, such an aid would require a large amount of clinical time to correctly “train”, as the user does not have direct control over what is optimal, and “unsharp inputs” are used in the training of the aid rather than precise and direct inputs.
Another example of a hearing aid that can receive the preferences of a hearing aid user and take these into consideration is described in U.S. Pat. No. 6,035,050. Such an aid requires the user to identify their current listening environment by selecting one of a limited number of listening situations on a remote unit. Should the user find that their current listening situation does not readily fall within those provided for selection, the benefit of such a aid becomes greatly reduced. Further to this, some listening situations on the remote unit, such as “at work”, may consist of a range of different acoustic environments and hence may not be acoustically well-defined. Thus, a neural network may not be able to reliably recognise such listening situations from an analysis of the microphone signal that, due to practical considerations such as memory limitations and the desirability of fast recognition of a change in listening situation, must be limited to a period of seconds or a few minutes. Therefore, after training it is possible that the neural network may misclassify a listening situation, especially in situations that are not acoustically well-defined, which can result in the application of amplification parameters that are unsuitable for the current listening situation. In such a device focus is placed on attempting to categorise specific acoustic environments, rather than measure the parameters of the environment and use these parameters directly in the processing scheme, as is the case with the present invention. Another disadvantage of this device is that a major lifestyle change, such as a new workplace that is acoustically different to the previous workplace, may require different amplification parameters and hence a return visit to the clinic.
Another example of a hearing aid that can receive the preferences of a hearing aid user and take these into consideration is described in U.S. Pat. No. 6,044,163. This document describes a hearing aid that is similar to the hearing aid described in U.S. Pat. 6,035,050. A major difference is that the neural network is not restricted to selection of sets of amplification parameters that are stored in a memory, but may directly set the value of individual amplification parameters. The disadvantages of the hearing aid described in this patent are similar to those of the hearing aid described in U.S. Pat. No. 6,035,050.
The present invention is adapted to providing users with an auditory prosthesis that is adaptable and can adjust its output to the user when that user is exposed to varying ambient acoustic environments.
Further, the present invention is adapted to providing users with an auditory prosthesis that can be ‘trained’ by the individual user to adapt its output to the user's preference in listening conditions encountered away from a traditional clinical situation, thereby reducing the clinical time required to fit an optimise a prosthesis to the individual user.
Any discussion of documents, acts, materials, devices, articles or the like which has been included in the present specification is solely for the purpose of providing a context for the present invention. It is not to be taken as an admission that any or all of these matters form part of the prior art base or were common general knowledge in the field relevant to the present invention as it existed before the priority date of each claim of this application.